FIG. 1 is a schematic axial view of a conventional third generation CT scanner which includes an x-ray source 12 and an x-ray detector system 14 secured to diametrically opposite sides of an annular shaped disk 16. The disk 16 is rotatably mounted within a gantry support (not shown), so that during a scan the disk 16 continuously rotates about a longitudinal z-axis while x-rays pass from the source 12 through an object, such as a patient 20, positioned on a patient table 56 within the opening of the disk 16. The z-axis is normal to the plane of the page in FIG. 1 and intersects the scanning plane at the mechanical center of rotation 18 of the disk 16. The mechanical center of rotation 18 of the disk corresponds to the “isocenter” of the reconstructed image.
In one conventional system, the detector system 14 includes an array of individual detectors 22 disposed in a single row in a shape of an arc having a center of curvature at the point 24, referred to as the “focal spot,” where the radiation emanates from the x-ray source 12. The source 12 and array of detectors 22 are positioned so that the x-ray paths between the source and each detector all lie in a “scanning plane” that is normal to the z-axis. Since the x-ray paths originate from what is substantially a point source and extend at different angles to the detectors, the diverging x-ray paths form a “fan beam” 26 that is incident on the detector array 14 in the form of a one-dimensional linear projection. The x-rays incident on a single detector at a measuring interval during a scan are commonly referred to as a “ray,” and each detector generates an output signal indicative of the intensity of its corresponding ray. The angle of a ray in space depends on the rotation angle of the disk and the location of the detector in the detector array. Since each ray is partially attenuated by all the mass in its path, the output signal generated by each detector is representative of the attenuation of all the mass disposed between that detector and the x-ray source, i.e., the attenuation of the mass lying in the detector's corresponding ray path. The x-ray intensity measured by each detector is converted by a logarithmic function to represent a line integral of the object's density, i.e., the projection value of the object along the x-ray path.
The output signals generated by the x-ray detectors are normally processed by a signal processing portion (not shown) of the CT system. The signal processing portion generally includes a data acquisition system (DAS) which filters the output signals generated by the x-ray detectors to improve their signal-to-noise ratio (SNR). The output signals generated by the DAS during a measuring interval are commonly referred to as a “projection,” “projection profile,” or “view” and the angular orientation of the disk 16, source 12 and detector system 14 corresponding to a particular projection profile is referred to as the “projection angle.”
If the detector array consists of N detectors, then N projection values are collected at each rotation angle. With the rays in a fan shape, these N projection values are collectively called a fan-beam projection profile of the object. The data of fan-beam projection profiles are often reordered or rebinned to become parallel-beam projection profiles. All rays in a parallel-beam profile have the same angle, called the parallel-beam projection view angle. The image of the object can be reconstructed from parallel-beam projection profiles over a view angle range of 180 degrees.
During a scan, the disk 16 rotates smoothly and continuously around the object being scanned, allowing the scanner 10 to generate a set of projections at a corresponding set of projection angles. In a conventional scan, the patient remains at the constant z-axis position during the scan. When obtaining multiple scans, the patient or the gantry is stepped along the longitudinal z-axis between scans. These processes are commonly referred to as “step-and-shoot” scanning or “constant-z-axis” (CZA) scanning. Using well-known algorithms, such as the inverse Radon transform, a tomogram may be generated from a set of projections that all share the same scanning plane normal to the z-axis. This common scanning plane is typically referred to as the “slice plane.”
A tomogram is a representation of the density of a two-dimensional slice along the slice plane of the object being scanned. The process of generating a tomogram from the projections is commonly referred to as “reconstruction,” since the tomogram may be thought of as being reconstructed from the projection data. The reconstruction process can include several steps including reordering to form parallel-beam data from the fan-beam data, convolution to deblur the data, and back projection in which image data for each image pixel is generated from the projection data. In CZA scanning, for a particular image slice, all the projections share a common scanning plane, so these projections may be applied directly for convolution and to the back projector for generation of a tomogram.
In some instances, for instance when the image consists of scans of the skull, the reconstructed images are further processed with a Beam Hardening Correction (“BHC”) operation, which is used to estimate regional tissue composition. This information is then used to adjust the predicted beam hardening transformation which is a function of tissue composition. The BHC is based on the assumption of a uniform average tissue composition. A monotonic transformation is predicted or measured based upon the average tissue composition. However, this procedure can result in certain artifacts being present in the corrected image in the form of a bleeding of the bone into the adjacent soft tissue which destroys the bone-brain interface in the image. These artifacts degrade the quality of the image and, in some cases, can render the image useless for certain diagnostic purposes.
Typically, a second-pass BHC operation or an Iterative Bone Correction (“IBC”) operation is performed on the BHC corrected image to compensate for the artifacts generated by the first-pass BHC. The effectiveness of the second pass correction depends on properties of the bone of the skull. It is more effective in images of adult skulls than in children's or infant skulls because of the increased density of the adult skull. Since the skull of an infant or child is less dense, the second-pass correction tends to overcorrect the image, resulting in a gap in the bone-brain interface in the image, which can equally render the image useless for diagnostic purposes. FIGS. 2A and 2B are reconstructed images of a CT scan of the skull of a two year old child. FIG. 2A shows a reconstructed image which has not had a second correction pass done. As can be seen, the first-pass correction resulted in a bleeding between the bone region of the image 12 and the brain region of the image 14, such that the bone-brain interface 16a is blurred and not easily discernable. FIG. 2B shows a reconstructed image which has had a second-pass correction performed. As can be seen, because the correction is performed at a gain which is used in bone correction operations of adult CT images, the lower density and area of the bone region 12 causes the image to be overcorrected, resulting in a gap in the bone-brain interface 16b between the bone region 12 and the brain region 14. In scans of infants (aged up to one year) the overcorrection is more pronounced, due to the low density and area of the bone in the skull.
In order to avoid the problems associated with the second pass correction in reconstructed scans, the second pass is not performed in younger patients. As stated above, in very young infants, the lack of the second pass correction does not adversely affect the scans because beam hardening is less likely to occur in the first pass correction operation. However, because of differences in the growth patterns of children, there is no clear age at which the second pass correction can be performed without resulting in some amount of overcorrection.